Optical coherence tomography (OCT) is an optical imaging method that has become an increasingly popular diagnostic tool in areas such as the biological, biomedical, medical screening, and vision-care. It utilizes low-coherence optical interferometry to enable non-invasive imaging of micron-scale microstructure inside biological tissue. In recent years, OCT has rapidly become competitive with radiography, ultrasound and magnetic resonance imaging in the biological and biomedical imaging communities due, in part, to its relatively low cost and high-resolution, in-vivo capabilities, as well as its lack of ionizing radiation. In the vision-care arena, for example, OCT is used to non-invasively image the human eye fundus, thereby facilitating diagnosis of retinal pathologies, such as macular degeneration, glaucoma, retinitous pigmentosa, and the like.
Early OCT systems were typically time-domain systems based on a relatively simple implementation of the classic Michelson interferometer. In such an interferometer, an input light signal is split into a reference arm and a sample arm, each of which reflects light back to a beam combiner that combines the two reflected light signals to generate an interference pattern that is then sampled by a detector. Only light that travels the same length in each of the reference arm and sample arm combine such that it results in a strong signal at the detector. In the reference arm, light is directed toward a movable reference mirror, which continuously reflects light back toward the detector. The length of the reference arm depends on the position of this mirror. In the sample arm, only light reflected by sub-surface features in the sample is returned toward the detector. The length of the sample arm, therefore, is based on the positions of features within the sample tissue. As a result, by scanning the reference mirror at a constant speed to change the length of the reference arm, the depth of each feature in the sample tissue is encoded in time by the position of the reference mirror, as represented in the interference pattern that is subsequently sampled by the detector.
Unfortunately, while early time-domain OCT techniques were promising, their complexity and time-intensive nature represented a significant limitation to their widespread adoption. Recent advances in OCT, however, have reduced the time necessary for a scan enabling the generation of high-resolution, cross-sectional (i.e., two-dimensional) images of sub-surface tissue microstructure.
Fourier-domain OCT (FD-OCT) introduced a scheme wherein a detector samples a resultant interferometric signal as a function of wavelength rather than position (i.e., time). This improved sampling scheme enables faster imaging with higher sensitivity. Typically, the broadband light is provided by either a broadband light source or a light source that sweeps through a range of optical frequencies (i.e., a swept source). In a swept-source system, each point of a two-dimensional area of a sample is sequentially illuminated with a monochromatic beam whose optical frequency is a function of time to generate an interferometric signal of intensity versus wavenumber, k (k is proportional to the inverse of wavelength). A mathematical algorithm, referred to as a Fourier transform, is then used to convert the interferometric signal to a plot of intensity versus depth in the tissue at the measured point. By virtue of its parallel nature, an FD-OCT system can generate a cross-sectional image of tissue features faster than early time-domain OCT systems.
The advantages of swept-source implementations over time-domain OCT include improved signal-to-noise ratio and faster scan rates. The wavelength-scanning speed of a swept-source FD-OCT system can exceed a wavelength sweep rate of 300 kHz using a research-grade Fourier-domain Mode-locking (FDML) laser. As a result, an FDML-based swept-source system is capable of generating a 100×100×512 pixel-volume dataset at 30 Hz (i.e., in real time). Unfortunately, this cross-sectional area is too small to have significant utility in many applications. Using a commercially available swept source, which has a typical sweep rate of approximately 50 kHz, an FD-OCT system could develop a 100×100×512 pixel-volume dataset at only 5 Hz, which is much slower than necessary for real-time imaging.
Spectral-domain OCT represents another approach for improving imaging speed. In a spectral-domain OCT system, light collected from all depths at each point of a two-dimensional area of a sample is dispersed across a linear array of detectors such that the information is sampled as a function of wavelength. Unfortunately, each point to be imaged must be integrated fully on the detector before another point can be imaged. As a result, the time required for detector integration and readout represents a significant bottleneck. Further, although no wavelength scanning is needed (in contrast to a swept-source system), the speed of the detectors used in spectral domain OCT systems are typically slower than the scan rate of the sources in swept-source systems. For example, typical commercial systems have equivalent line rates of 29 kHz, which corresponds to a maximum image rate of 2.9 Hz for a 100×100×512 volume dataset. Recent improvements in commercially available linear detectors have improved the potential line rate to 92 kHz; however, this corresponds to a maximum image rate of only approximately 9.2 Hz, still woefully inadequate for real-time imaging applications.
Spectral encoded endoscopy (SEE) is a particular implementation of spectral-domain OCT that uses a single optical fiber and miniature diffractive optics to obtain endoscopic images through small diameter probes. Using spectral-domain interferometry, SEE is furthermore capable of three-dimensional volume imaging at video rates. In a SEE system, light from a swept source or broadband source is dispersed linearly along N image points of a two-dimensional area of a sample. As a result, each of the image points is illuminated by 1/N of the wavelength range of the source light. Although SEE enables high-speed volumetric imaging, the resolution of the image suffers since depth resolution depends on the total wavelength range used to interrogate an image point.
Even with the advances in OCT modalities, real-time imaging of three-dimensional volumes of reasonable size and resolution still eludes the OCT community.